In the present specification, reference is made to the following publications illustrating prior art of conventional multi-channel coils for magnetic resonance (MR) applications, which are incorporated by reference in their entirety herein:    [1] H. Hetherington, N. Avdievich, A. Kuznetsov, and J. Pan: “RF Shimming for Spectroscopic Localization in the Human Brain at 7 T” in “Magn. Reson. Med.” 63:9-19 (2010);    [2] U. Katscher, P. Bornert, C. Leussler, J. S. van den Brink: “Transmit SENSE” in “Magn. Reson. Med.” 2003; 49: 144-150;    [3] G. Adriany, J. Ritter, T. Vaughan, K. Ugurbil, P.-F. Van de Moortele: “Experimental Verification of Enhanced B1 Shim Performance with a Z-Encoding RF Coil Array at 7 Tesla” in “Proceedings of the 18th Annual Meeting” 2010, p. 3831, May 2010;    [4] K. M. Gilbert, A. T. Curtis, J. S. Gati, L. M. Klassen, R. S. Menon: “A radiofrequency coil to facilitate B(1) (+) shimming and parallel imaging acceleration in three dimensions at 7 T” in “NMR Biomed” Dec. 8, 2010 [Epub ahead of print];    [5] G. C. Wiggins, A. Mareyam, K. Setsompop, V. Alagappan, A. Potthast, L. L. Wald: “A close-fitting 7 Tesla 8 channel transmit/receive helmet array with dodecahedral symmetry and B1 variation along” in “Proceedings of the 16th Annual Meeting of ISMRM” Toronto, Canada, p. 148, May 2008; and    [6] N. Avdievich, J. Pan, and H. Hetherington: “Improved longitudinal coverage for human brain at 7 T: A 16 Element Transceiver Array” in “Proceedings of the 19th Annual Meeting of ISMRM” Montreal, Canada, p. 328, May 2011;    [7] WO 2011/029452; and    [8] N. I. Avdievich, J. W. Pan, J. M. Baehring, D. D. Spencer and H. P. Hetherington: “Short Echo Spectroscopic Imaging of the Human Brain at 7 T Using Transceiver Arrays” in “Magn. Reson. Med.” 62:17-25 (2009).
For magnetic resonance imaging, especially at high level of a static magnetic field (oriented in an axial direction), use of a multi-channel transmit coil for producing the desired RF field becomes increasingly important, because such a coil provides additional degrees of freedom that enable static RF shimming (adjustment of amplitude and phase of excitation signals to obtain better homogeneity).
The coil can be excited: 1) by a single-channel transmit power source connected to multi-channel power splitter followed by a phase shifter, such that the relationship between RF amplitude and phase for each channel remains the same for any slice excitation; or 2) by a multi-channel power transmitter unit, with which the homogeneity of each slice can be further improved (on a slice-by-slice basis) by using different RF shim settings for each slice excited [1].
It has been shown in practice that when the electromagnetic field wavelength is comparable with the dimensions of the volume of interest (VOI) (for instance, a human head at a static magnetic field of 7 T and above, thus a Larmor frequency of more than 300 MHz), a single row of coil elements placed around the azimuth of a cylindrical coil former cannot produce a highly uniform RF field within the entire VOI, no matter how the individual coils are driven using the RF shim principle. With such a coil geometry and slice-by-slice optimization, it is possible to achieve reasonable RF homogeneity within transverse slices, but this cannot be achieved with coronal or sagittal slices.
Use of transmit SENSE [2] (dynamic RF shimming—simultaneous and time dependent adjustment of amplitude/phase of excitation signals and scanner gradient field) has been proposed to improve homogeneity in any given slice. This approach allows very high homogeneity to be obtained for any slice within the VOI, at the expense of the need to repeat the laborious preparation procedure for each subject (and even each position of a given subject within a coil).
The preparation procedures both for static and dynamic RF shimming include the rather time consuming step of RF field mapping (for all slices excited), followed by 1) calculation/optimization of required RF shim sets for static RF shim; and 2) optimization of the RF pulse sequence (amplitude and phase at each excitation time step) for each individual transmit channel and scanner gradient sequence for RF dynamic shim. In most cases, the results (the RF shim sets or RF pulse sequence) obtained are dependent on the coil loading, in particular varying with human subject. In general, static and dynamic RF shimming requires correct coil tuning, matching and decoupling, which are also, in most cases, load dependent.
The maximum value within the load of the local specific absorption rate (SAR) and the global SAR significantly depends on the specific combination of RF amplitudes and phases of individual transmit channels. For this reason, use of either static or dynamic RF shimming requires a reliable real-time SAR monitoring system. The hardware infrastructure required for this is absent from present-day commercially available MRI scanners, resulting in their sub-optimal usage, especially for clinical applications, in which the allowed SAR values are much smaller than strictly required for patient safety.
Because implementation of static RF shimming is much less complicated than dynamic RF shimming, the strategy has been proposed of using multiple rows of coils to improve the RF homogeneity in the axial direction while using only static RF shimming [3, 4, 5, 6]. These studies have shown that this indeed improves useful axial coverage, but results in additional complexity in coil design and fabrication, because all elements must be separately adjusted and matched, as well as decoupled from all adjacent elements.
The most common RF coil tuning, matching and decoupling procedure is the minimization of each element's reflection coefficient (Sxx) and its coupling (Sxy) with adjacent elements. This usually entails that all Sxx and Sxy, values approach their minimum at a frequency equal to the MRI Larmor frequency (fres). This condition is relatively easy to achieve for coils with a moderate number of elements, but for a multi-row array, in which there are at least four adjacent elements (depending on the specific spatial design), the full procedure is challenging and lengthy [4].
If after tuning, matching and decoupling the Sxy values are not small enough, the power reflected by the entire coil (Pref_coil) can amount to a substantial fraction of the entire transmit power (Ptransmit). This can result in severe degradation of the coil performance estimated as B1+VOI/√Ptransmit, where B1+VOI is B1+ averaged over the VOI.
Direct minimization of the total reflected power Pref_coil, with the omission of any dedicated decoupling network, has been shown to provide a simple method for achieving the maximum near-field magnetic field generated per unit delivered power, for most single-row closed loop multi-channel coils, independently of the design of the array radiative elements (loop, shielded microstrip, stripline). However, this approach fails for the multi-row array. While the method easily handles the coupling between azimuthally adjacent coil elements (and in most cases benefits from it), there are no degrees of freedom available to compensate for axial coupling.